Magnetic resonance imaging (MRI) is a technique that is capable of providing three-dimensional imaging of an object. A conventional MRI system typically includes a main or primary magnet that provides the main static magnetic field Bo, magnetic field gradient coils and radio frequency (RF) coils, which are used for spatial encoding, exciting and detecting the nuclei for imaging. Typically, the main magnet is designed to provide a homogeneous magnetic field in an internal region within the main magnet, for example, in the air space of a large central bore of a solenoid or in the air gap between the magnetic pole plates of a C-type magnet. The patient or object to be imaged is positioned in the homogeneous field region located in such air space. The gradient field and the RF coils are typically located external to the patient or object to be imaged and inside the geometry of the main or primary magnet(s) surrounding the air space. There is shown in U.S. Pat. Nos. 4,689,563; 4,968,937 and 5,990,681, the teachings of which are incorporated herein by reference, some exemplary MRI systems.
In MRI, the uniform magnetic field Bo generated by the main magnet is applied to an imaged object by convention along the Z-axis of a Cartesian coordinate system, the origin of which is within the imaged object. The uniform magnetic field Bo being applied has the effect of aligning the magnetization arising from the nuclei of the atoms comprising the imaged object, along the Z-axis, such nuclei possess a nuclear magnetization due to their having an odd number of protons or neutrons. In response to RF magnetic field pulses of the proper frequency, with field direction orientated within the XY plane, the nuclei resonate at their Larmor frequencies, ω=γBo where γ is called the gyromagnetic ratio. In a typical planar imaging sequence, the RF signal centered about the desired Larmor frequency is applied to the imaged object at the same time a magnetic field gradient Gx is being applied along the Z-axis. This gradient field Gx causes only the nuclei in a slice of limited thickness through the object perpendicular to the Z-axis, to satisfy the resonant condition and thus be excited into resonance.
After excitation of the nuclei in the slice, magnetic field gradients are applied along the X- and Y-axes respectively. The gradient Gx along the X-axis causes the nuclei to precess at different frequencies depending on their position along the X-axis, that is, Gx spatially encodes the precessing nuclei by frequency. Thus, this gradient is often referred to as a frequency encoding or read-out gradient. The Y-axis gradient Gy is incremented through a series of values and encodes the Y position into the rate of change of the phase of the precessing nuclei as a function of gradient amplitude, a process typically referred to as phase encoding.
The quality of the image produced by the MRI techniques is dependent, in part, upon the strength of the magnetic resonance (MR) signal received from the precessing nuclei. For this reason an independent RF coil is often placed in close proximity to the region of interest of the imaged object, more particularly on the surface of the imaged object, in order to improve the strength of the received signal. Such RF coils are sometimes referred to as local coils or surface coils.
There is described in U.S. Pat. No. 4,825,162 a surface coil(s) for use in MRI/NMRI imaging and methods related thereto. In the preferred embodiment of that invention, each surface coil is connected to the input of an associated one of a like plurality of low-input-impedance preamplifiers, which minimize the interaction between any surface coil and any other surface coils not immediately adjacent thereto. These surface coils can have square, circular and the like geometries. This yields an array of a plurality of closely spaced surface coils, each positioned so as to have substantially no interaction with all adjacent surface coils. A different MR response signal is received at each different one of the surface coils from an associated portion of the sample enclosed within the imaging volume defined by the array. Each different MR response signal is used to construct a different one of plurality of different images from each surface coil. These images are then being combined, on a point-by-point basis to produce a single composite MR image of a total sample portion comprised of the MR response signals from the entire array of surface coils.
The use of a phased array of RF coils or surface coils with a tuned and matched circuit including low impedance pre-amplifiers have been used to de-couple adjacent loops as a mechanism for improving the signal-to-noise ratio (SNR) and field of view (FOV). In this regard, it should be understood that the term “coupling” refers to the coupling of a signal (e.g., MR signal and/or noise signal and/or RF excite signal) in one coil to an adjacent coil(s), such that the signal being outputted by the adjacent coil is a combination of the signal detected by the adjacent coil and the coupled signal. Consequently, the image from the adjacent coil may be distorted or the SNR degraded to some degree. Although overlapping adjacent coil(s) and using low impedance pre-amplifiers have been effective in minimizing decoupling artifacts and SNR degradation, such circuitry becomes less effective as the number of coils and/or the coil density is increased. In particular, as the spacing between adjacent coils and between adjacent portions of a coil is decreased signal coupling effects increase and become less manageable by the various measures deployed to counter such effects.
Although there are a variety of spatial encoding methodologies or techniques being implemented, a popular method used in commercial MRI scanners is two-dimensional Fourier transform (2DFT) encoding in which a two-dimensional spatial plane (e.g., XY plane) is first selectively excited then encoded with both frequency and phase of the MR signals. Typically during one data acquisition, only a one dimensional time-domain signal is obtained from the plane and the 2DFT encoding requires repeating the data acquisitions sequentially to achieve a pseudo second dimension of the time domain signals. This second dimension of spatial information is encoded by repeating the data acquisition with different phase encoding gradient strengths (i.e., varying Gy to create the pseudo-time dimension). The spatial resolution in the Y-direction or axis generally corresponds to the FOV divided by the number of different gradient strengths used (e.g., 128, 256). The number of different gradient phase-encoding steps also generally corresponds to the number of pixels in the Y-direction. The spatial information in the X-direction is encoded in the time-domain signal that is acquired in the presence of the constant frequency-encoding, or readout, gradient. Because the spatial resolution depends in part on the number of repetitions and therefore the number of phase-encoding steps, and because the repetition rate is limited by the MR relaxation times; a higher resolution image generally takes a longer time.
Two recent methods, the Simultaneous Acquisition of Spatial Harmonics (SMASH) imaging in the time domain and the Sensitivity Encoded (SENSE) imaging in the frequency domain, replace such sequential data acquisition by a partially parallel process by using a multi-channel phased array detector system, thereby reducing the scan time as compared to the sequential data acquisition technique. In these two techniques, it is recognized that the data lost by sampling below the Nyquist sampling rate can be recovered using the spatial information intrinsically contained in the sensitivity profiles of the individual detectors that comprise the phased array.
Thus, the time domain method (SMASH) recognizes the equivalence between phase-encoding with MRI gradients and the composite spatial harmonics that are present in the detector sensitivity profiles. It uses a numerical fitting routine to recreate a number (i.e., usually a limited number) of phase-encoding steps from the profiles, and thus permits reductions in scan time by allowing the number of conventional phase encoding steps to be reduced by the number of steps that can be so recreated. Although this numerical approach was instrumental in demonstrating the original SMASH concept, it did not recognize the underlying analytical relationship between the weighting factors for the composite harmonics, the FOV, the spacing of the detectors, the harmonic orders, and the sensitivity profiles of the detector coils, which was addressed subsequently by the Analytic SMASH Procedure (ASP), described in Lee et al, Magn Reson Med 2000; 43: 716–725. See also U.S. Ser. No. 09/728,948 entitled “Method for Parallel Spatial Encoded MRI and Apparatus, Systems and other Methods Related Thereto, the teachings of which are incorporated herein by reference, owned by the Assignee of the present invention.
A problem with all of these methods (SMASH, SENSE, ASP, etc.) is that conventional MRI phased array coils are limited in the number of coils that can be deployed, by the limitations imposed by both their loop structure and the de-coupling requirements for the mutual induction between the elements. Because the number of coils in the phased array correlates to the maximum factor for reducing the number of phase encoding steps, existing phased array designs significantly limit the potential for parallel spatial encoding.
There is found in U.S. Ser. No. 09/822,771 entitled “APPARATUS FOR MAGNETIC RESONANCE IMAGING HAVING A PLANAR STRIP ARRAY ANTENNA INCLUDING SYSTEMS AND METHODS RELATED THERETO” (owned by the assignee of the subject patent application) a new type of phased-array detector, a planar strip array (PSA). The PSA includes an array of parallel conductive strips that are disposed on one surface of a substrate and on the other opposing surface thereof is disposed a ground plane. An overlay is applied so as to cover the conductors disposed on the substrate. Each of the conductor strips has a length that is set so as to substantially reduce the coupling of a signal in one of the conductors to an adjacent conductor. More particularly, the conductor length is set so as to be equal to be about nλ/4, where n is an integer ≧1 and λ is the electromagnetic (EM) wavelength of the signal to be detected from a sample placed on top. For MRI applications, λ is the wavelength in the detector corresponding to the NMR resonance frequency for the nuclei in the sample being subjected to a given magnetic field strength by the main or primary magnetic coils. The materials comprising the overlay and the substrate are comprised of a material whose relative permittivity material is chosen to result in an EM wavelength such that the quarter wavelength (or nλ/4) condition produces a detector whose length is reasonable relative to the sample or image FOV.
Under these conditions, the behavior of the PSA is such that it provides decoupling of adjacent and neighboring elements or conductive strips. Unlike loop coils, the conductive strips are inherently decoupled to a certain extent regardless of their spacing from each other, and are broadband decoupled when the spacing “s” and the total thickness “h” of the substrate and overlay approximately satisfy the relation s≧2.5h. Because the electrical field is concentrated between the conductive strips and the ground plane, losses associated with the electrical field are minimized. Thus, the PSA structure is capable of accommodating a large number of detectors for simultaneously acquiring near-field MRI signals without interference between each other, and is ideally suited to conventional phased-array MRI or parallel sensitivity encoded MRI. Although the PSA has made important and significant advances with regards to decoupling of adjacent conductive strips or elements, there is a continuing need to improve upon the capabilities of a parallel conductive strip array. More particularly, there also is a need to expand upon the capabilities of the PSA teachings so a parallel conductive strip array (e.g., configuration and size) can be optimized for various anatomies and organs of interest, such as can be done with conventional loop arrays. Also, it would be desirable to expand upon the capabilities of the PSA teachings so a parallel conductive strip array has a non-planar, cylindrical geometry for applications for the head, limbs or other generally cylindrically shaped objects as well as having a curved geometry so that the array conforms to the general contour of the object to be sampled, for example a body.
It thus would be desirable to provide an array antenna or RF MR signal detector, as well as systems and methods embodying such an antenna/detector, in which adjacent antenna or conducting elements are decoupled and which array antenna is tunable to any of a number of MRI frequencies. It also would be desirable to provide such an array antenna or RF MR signal detector, as well as systems and methods embodying such an antenna/detector, where the length of the conducting elements can be selected based on MR performance criteria specific to the geometry of the sample and region of interest or organs to be imaged. More particularly, it would be desirable to provide such an array antenna or RF MR signal detector, as well as systems and methods embodying such an antenna/detector, where the length of the conducting elements can be adjusted to allow one to optimize SNR performance for a specific depth or range of depths in the body or sample being scanned. In addition, it would be desirable to provide such an array antenna or RF MR signal detector in which there is a reduced sensitivity variation as a function of position along the conducting elements. Further, it would be particularly desirable to provide such an array antenna or RF MR signal detector, as well as systems and methods embodying such an antenna/detector, that embodies a non-planar geometrical configuration, more specifically a cylindrical geometrical configuration. In addition, it would be desirable to provide such an array antenna/detector, system and method in which the array antenna includes a plurality or more, more particularly a multiplicity, of elements for detecting MR signals, such as 4 or more, 16 or more, 32 or more or 64 or more elements so as to significantly reduce scan time, particularly when the array is in conjunction with parallel sensitivity encoding methods, such as SMASH, SENSE, ASP and the like. Such an array antenna/MR signal detector preferably would be simple in construction as compared to prior art loop phased array antennas and/or MR signal detectors.